The concentration of a medically interesting analyte in a biological fluid is often determined by means of an electrochemical test. For example, a highly specific reaction between the analyte and a test reagent, generally an enzyme, may be carried out which causes a change of the oxidation stage of the enzyme. The degree to which this change takes place depends on the concentration of the analyte. The degree of the change is detected by means of a charge transfer to a solid electrode. The rate of charge transfer at the electrode is detected by applying a DC-voltage thereto and measuring the resulting DC-current. The concentration of the analyte is derived from this current measurement.
A typical reaction sequence of such an electrochemical test is shown in FIG. 1. An enzyme catalyzes an oxidation of an analyte A thereby forming an analyte product P. Simultaneously the enzyme is reduced from its oxidized form Eox to its reduced form Ered. By this highly specific enzymatic reaction an electron is transferred from the analyte to the enzyme. The enzyme cycles back to the oxidized form thereby transferring an electron to a mediator which is transferred from its oxidized form Mox to its reduced form Mred. The mediator is reoxidized by application of a DC-voltage, thereby transferring an electron to a solid state electrode EL.
Examples of the use of this test principle include glucose tests which can be used to monitor the blood sugar status of diabetics. This monitoring is of utmost importance both medically and commercially, because the health status of diabetics highly depends on reliable, easily generated data on their blood glucose level. The enzymes used in glucose testing are generally glucose oxidase or glucose dehydrogenase. Prominent mediators are ferrocen, ferrocyanide and phenylendiamine. Other enzymes and mediators are also available in the art.
As described herein, the present invention is useful with enzymatic amperometric tests, such as for glucose. It can, however, be useful with any amperometric test which includes an overall reaction sequence comprising the following redox reactions:
an analyte-specific redox reaction in a homogeneous (hydrated) phase by which a product is formed which can, with application of a suitable voltage, be oxidized or reduced at a solid electrode (also referred to as a “forward reaction”);
a heterogeneous electrode reaction by which an electron is transferred from the product of the analyte-specific reaction to the electrode or from the electrode to the product (also referred to as a “reverse reaction” or “electrode reaction”).
Clearly, each of these reactions may comprise a plurality of reaction stages. For example, the electrochemical test sequence that includes the formation of Mred via the enzyme, as shown in FIG. 1, is a two-stage analyte-specific reaction. The mediator is, however, only needed if the direct charge transfer of the enzyme to the solid state electrode is too slow. Therefore, depending on the type of enzyme and electrode used, tests are also possible in which the mediator is omitted. On the other hand, the present invention may be used in electrochemical tests comprising several reaction stages.
Electrochemical tests are produced in numerous different formats including electrodes suitable for immersion into the sample liquid. The present invention is applicable to different formats. In one embodiment, use of the invention is directed to biosensor systems comprising disposable analysis elements and an evaluation instrument which is specifically adapted for evaluation thereof. Normally, the components of the system are developed and supplied by the same manufacturer. The analysis elements contain the reagent system required for a particular test (designated here as “analyte reactant”) and at least two electrodes. The evaluation instrument of the system generally has a holder for receiving a disposable analysis element (also referred to herein as a “biosensor”). When the biosensor is plugged into the holder, an electrical contact is established between the electrodes of the biosensor and the instrument electronics.
In various embodiments of a biosensor, the analyte reactant may be contained in a porous matrix such as a paper or porous plastic element which is in contact with the electrodes and to which the sample is applied. In recent years, an alternative design of analysis elements has become increasingly important in which the analyte reactant and the electrodes are contained in a capillary space. The sample liquid is applied to an opening of the capillary space and drawn into the space by capillary action. The analyte reactant is at least partially hydrated by the sample liquid drawn into the space, thereby forming a reaction layer which is in contact with the electrodes provided with the biosensor for measuring the analyte detection current which corresponds to the analyte concentration in the sample liquid. Various embodiments of the present invention described herein relate generally to biosensor systems comprising such a “capillary biosensor”. Nevertheless, a person of ordinary skill in the art of biosensor systems will understand and appreciate alternative embodiments of the invention including use of biosensors other than capillary biosensors, and such alternatives are intended to be within the scope of the claims appended hereto, except as may be otherwise recited therein.
More details about such biosensor systems can be taken from the appropriate literature. In particular, a capillary biosensor system is described in WO 99/32881 and this document contains an extensive list of earlier publications from which a substantial amount of additional technical information can be taken. The disclosure of this document and of the documents listed therein are incorporated herein by reference.
FIG. 2 is taken from U.S. Pat. No. 5,243,516 and shows a typical example of the timing sequence in prior art electrochemical tests. After an analysis element has been plugged into the evaluation instrument, a sample detection voltage is applied to the electrodes. Substantially no current flows as long as there is no sample in a sample well of the analysis element bridging the electrodes. However, as soon as the sample well is sufficiently filled, a current spike CS that exceeds a given threshold is sensed, indicating that a drop of sample has been dosed into the sample well of the analysis element and the resulting reaction layer formed by the sample and the analyte reactant bridges the electrodes. This point of time is designated “dose detect” (DD). After current spike CS has been sensed, the sample detect voltage is removed from the electrodes and the analyte-specific forward reaction takes place during an incubation period IP. After the incubation period IP, an analysis voltage suitable for the electrode reaction (reverse reaction) is applied to the electrodes. In one embodiment, the analysis voltage is larger than the sample detection voltage. In other embodiments, the analysis voltage is substantially the same as the sample detection voltage. The differences in selecting the value of the analysis voltage as compared to the sample detection voltage is generally understood and appreciated in the art to those of ordinary skill, for example as described in U.S. Pat. No. 5,243,516. FIG. 2 shows typical shapes of the resulting functional relationship I(t) of the current I versus time for four different values of the glucose concentration as indicated in the Figure. These curves of the DC-current versus time are hereafter designated “I(t)-traces”.
If all required test conditions are met, the shape of the l(t)-traces corresponds (after a surge time ST in which the I(t)-trace is substantially influenced by the particulars of the measurement electronics) to a characteristic function proportional to 1/√t. A deviation from this “Cottrell current curve” indicates deviations from the required test conditions. U.S. Pat. No. 5,243,516 proposes to make a plurality of current measurements at a plurality of measurement times during the period in which the reverse reaction takes place and to use a simple mathematical method to control whether the I(t)-traces as shown in FIG. 2 are in agreement with the Cottrell current. If this is not the case a malfunction of the system can be assumed and indicated.
Problems with accuracy of biosensor system measurement results are caused by sources of error which have effects on the biosensor output, such as temperature and other interferents. An example of an interferent factor is the concentration of red blood cells, i.e. hematocrit, in whole blood. This problem is discussed in the publication:                Tang et al.: “Effects of Different Hematocrit Levels on Glucose Measurement with Handheld Meters for Point-of-Care Testing”, Arch Pathol Lab Med, 2000, pp. 1135 to 1140.        
The authors point out that, even with the latest technology of biosensors, large errors of the glucose value on the order of 20 to 30% are caused by variations of the hematocrit which may occur in practice. A number of possible mechanisms by which these errors may be caused are mentioned. It is noted that solutions to this problem are badly needed but no means for compensation of the hematocrit error are described by those authors.
Another important source of error are temperature variations. Amperometric test results generally show a strong dependence on the temperature of the reaction layer. Therefore, in some biosensor systems, the temperature of the reaction layer is carefully controlled to a fixed value. In other systems the temperature is measured and a correction calculation is performed to compensate for temperature variations.
According to the above-mentioned WO 99/32881, an AC-measurement can be made to achieve a correction for the combined effect of sample temperature and hematocrit. To this end, an AC-voltage in the frequency range between about 2 kHz and about 10 kHz is applied to the electrodes, and the real and imaginary components of the impedance of the biosensor-sample system are determined. The magnitude and the phase angle of the impedance are calculated therefrom and a look-up table stored in the instrument is consulted for a correction factor. This correction factor is applied to conventionally determined glucose values, thereby deriving a corrected glucose concentration.